Medical imaging is one of the most useful diagnostic tools available in modern medicine. Medical imaging allows medical personnel to non-intrusively look into a living body in order to detect and assess many types of injuries, diseases, conditions, etc. Medical imaging allows doctors and technicians to more easily and correctly make a diagnosis, decide on a treatment, prescribe medication, perform surgery or other treatments, etc.
There are medical imaging processes of many types and for many different purposes, situations, or uses. They commonly share the ability to create an image of a bodily region of a patient, and can do so non-invasively. Examples of some common medical imaging types are nuclear medical (NM) imaging such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), electron-beam X-ray computed tomography (CT), magnetic resonance imaging (MRI), and ultrasound (US). Using these or other imaging types and associated machines, an image or series of images may be captured. Other devices may then be used to process the image in some fashion. Finally, a doctor or technician may read the image in order to provide a diagnosis. Thus, systems and methods which enhance the accuracy of the images can be beneficial in describing and treating medical conditions.
Positron emission tomography (PET) is a nuclear medicine imaging technique that produces a three-dimensional image or picture of functional processes in the body. The system detects pairs of gamma rays emitted indirectly by a positron-emitting radionuclide (tracer), which is introduced into the body on a biologically active molecule. Three-dimensional images of tracer concentration within the body are then constructed by computer analysis. Data collection in PET can involve the use of scintillation detectors. A scintillation detector or scintillation counter is obtained when a scintillator is coupled to an electronic light sensor.
Photosensors can include an array of independent Geiger-mode avalanche photodiode (APD) cells, each with an integrated quenching resistor. When an individual APD absorbs one or more photons, it may go into avalanche. The avalanche is quenched as current flows through the quenching resistor, producing a bias voltage drop on the diode. Since all of the APDs are connected to a common electrode, if the SiPM is hit by a pulse of light, the charge dumped onto the electrode will be proportional to the number of APDs that fire, and, therefore, proportional to the number of incident photons.
Scintillation detection is one application of SiPMs in which pulses of light, often containing large numbers of photons, must be detected. For SiPMs, however, there is a trade-off between photon detection efficiency (PDE) and linearity. For a fixed SiPM area and fixed dead-space between individual elements, as the number of APD cells in the array is decreased, the geometric efficiency increases, resulting in higher PDE. For high intensity light pulses (i.e. conversion of high energy gamma rays in the scintillator) the number of APDs that absorb multiple photons also increases as the number of APDs is decreased. Since the charge produced by a single APD in Geiger mode is independent of the number of photons absorbed, the response of the SiPM becomes more non-linear.
The impact of non-linearity on the average signal level (i.e. the peak positions in a pulse-height spectrum) can be corrected by proper calibration. In additional to changing the peak positions, non-linearity affects the energy resolution of a scintillation detection system. If the non-linearity becomes severe enough, it will significantly degrade the measured energy resolution of the system, which can result in improper imaging.
Thus, there exists a need in the art to correct for the above described non-linearities in order to facilitate imaging accuracy.